1. Field of the Invention
This invention relates to a self-regulated implantable pump for the delivery of insulin to diabetic mammals, particularly humans, in response to their insulin needs. Specifically, it relates to a mechanochemical pump which, when implanted in a mammal, delivers insulin by hydraulic force in direct and proportionate response to an increase in the blood glucose level of the mammal in which it is implanted.
Approximately two million Americans are insulin dependent diabetics. After heart disease and cancer, diabetes is the third leading cause of death in the United States. It is also the leading cause of blindness in the U.S., and a major cause of neuropathy, nephropathy and atherosclerosis. Before the isolation of insulin in 1921 by Banting and Best, insulin dependent diabetics usually died within four years of onset due to the inability to efficiently utilize glucose as an energy source, and the ketoacidosis which arises from breaking down fat for energy. The introduction of insulin therapy has substantially eliminated the acutely lethal aspects of insulin dependent diabetes. Mortality, however, remains high, due primarily to the longer term impact of the degenerative conditions listed above.
2. Description of Related Art
Presently, the most common route of insulin delivery is by subcutaneous injection. This procedure is inconvenient to the patient, and is probably not optimal from a therapeutic standpoint. Blindness, neuropathy and nephropathy are caused by the thickening of capillary basement membranes in the corresponding organs. This, as well as atherosclerosis, stems from a general hyperglycemia which results from the fact that the release of subcutaneously injected insulin does not mimic the release pattern of insulin from a healthy pancreas. Diabetics can bring their blood sugar level down into the normal range by careful control of diet. However, in this case there is always the danger that blood sugar will drop below acceptable levels (hypoglycemia), which is a dangerous condition.
The normal pancreas releases insulin at a rate determined by blood sugar level. By this method the nondiabetic body exercises closed loop control of blood sugar. In contrast, subcutaneous injection is clearly an open loop strategy, which explains the suboptimal glycemia control that is achieved. As a result, diabetics are usually hyperglycemic.
It has been hypothesized that the long term degenerative complications of diabetes might be averted if blood glucose levels could be reduced. This could be accomplished by an implantable, self-regulating insulin delivery system. In addition, such a system would eliminate the inconvenience of regular injections, albeit at the cost of requiring occasional surgery. The present goal is to develop an implantable system that could provide self-regulated delivery for a period of six months to one year.
Several research groups have developed insulin delivery systems that attempt to mimic the closed loop aspect of the normal pancreas. These can be classified as electromechanical and chemical. The electromechanical devices generally require an electrochemical glucose sensor which signals to an electromechanical actuator that pumps glucose. In chemical systems, no electrical component exists, and the sensing of glucose and transduction to insulin release rate are coupled solely by chemical processes.
To date the most fully developed system for closed-loop insulin delivery is the electromechanical "artificial beta cell," developed by Albisser and colleagues [Proc. IEEE 67, 1308 (1979)]. This is an experimental bedside system. The patient's blood is sampled by a catheter and analyzed continuously for glucose. Glucose levels are fed into a computer which then determines the insulin delivery rate. Insulin delivery is actuated by a peristaltic pump feeding a second catheter to the patient. This system is mainly an experimental tool, since the collective bulk of all components, plus the need for a cutaneous catheter, preclude its use with ambulatory patients.
Attempts to miniaturize this and other electromechanical systems for in vivo implantation have been thwarted by the unavailability of a chronically implantable glucose sensor. It is not expected that this problem will be solved in the near future. Even if miniaturization were possible, certain drawbacks would remain for any electromechanical system. A miniaturized artificial beta cell would be either wearable or implanted. Wearable systems, although doubtless an improvement over injections, necessarily involve a cutaneous junction which is inconvenient, and a potential locus for infection if great care is not taken. Implantable electromechanical systems also require a power source, which takes up considerable volume, plus many moving parts, all of which are subject to failure.
The potential problems with electromechanical insulin delivery have led several groups of researchers to consider chemical means to provide closed-loop insulin delivery. Kim and colleagues [J. Controlled Release 1, 57 (1984), J. Controlled Release 1, 67, (1984) and J. Controlled Release 2, 143 (1985)] have designed a system that takes advantage of the fact that glycosylated insulin bound to the plant lectin concanavalin A (ConA) is displaced by glucose. A device was developed consisting of a macroporous, hydrophilic membrane surrounding a solution of ConA/glycosylated insulin. When blood glucose level rises, glucose crosses the membrane and displaces the glycosylated insulin from the ConA. The released insulin diffuses through the membrane into the blood. However, a potential problem with this system is the use of glycosylated insulin and ConA, each of which may induce an immunogenic response.
Another idea has been pursued by Ghodsian, et al. [Proc. Natl. Acad. Sci. 85, 2403 (1988). A mixture of the immobilized enzymes glucose oxidase (GluOx) and catalase (Cat), and a chemically derivatized form of insulin (trilysil insulin) was incorporated into a porous polymer matrix. The enzymes convert glucose oxidase into gluconic acid, causing a local drop in pH within the pores of the polymer matrix. Since the solubility of trilysil insulin is minimal at neutral pH (7.4), and increases rapidly with decreasing pH, an increase in insulin solubility and therefore release is the ultimate result of an increase in glucose concentration. In principle this system can provide closed loop control. At present there appear to be two disadvantages to this approach, however. First, the suitability of trilysil insulin for long term use has not been established. Unmodified human, porcine or bovine insulins cannot be used, since their solubility-pH profiles are not correct. Second, the response time becomes slower with each burst of insulin. This is because release occurs at a "moving front" which separates solid insulin from pore water. With time, this front continuously recedes into the matrix, and more time is required for glucose and H+ to diffuse in the pores to the insulin, and for insulin to diffuse out of the pores.
A third polymer-based chemically selfregulating system has been pursued by Heller [Proc. 13th Intl. Symp. Controlled Release of Bioactive Materials, p. 35 (1986)]. In this system, insulin is imbedded into a hydrophobic polymer which undergoes surface erosion. As the polymer erodes, insulin at and just below the surface is released. The polymer is chosen such that the erosion rate is small at neutral pH but increases as pH decreases. Again, change in pH is triggered by the GluOx/Cat reaction.
Yet another approach is being pursued by Horbett, et al. [J. Biomed. Mater. Res. 19, 1117 (1985) and J. Controlled Release 6, 267 (1987)], and Ishihara, et al. [Polymer J. 16, 625 (1984)]. These research groups have developed cross-linked hydrogel membranes whose permeability to insulin is pH sensitive. Permeability change is due to swelling of the polymer and "opening of pores." The swelling in turn is due to protonation of amine groups residing on the sidechains of the polymer as a result of the pH-lowering effect of the glucose/glucose oxidase reaction. Practical implementation of these membranes would involve an aqueous insulin reservoir surrounded by the permeable membranes. However, with this system, any insulin degradation and aggregation which occur over time could lead to inactivation of the drug and clogging of the membrane.
Thus, while the concept of utilizing the reaction of blood glucose with glucose oxidase to trigger a pH-responsive delivery of insulin is known, the art available insulin delivery mechanisms which utilize this reaction require the transport of insulin from and through the very matrices and membranes in which the glucose/glucose oxidase reaction takes place. As a result, the self-regulating delivery properties of these matrices and membranes are highly susceptible to deposits of insulin degradation products, and in the case of the matrix, suffer from delivery kinetics which slow with time due to the geometry of the matrix and the receding insulin front.